Over 100,000 people in the United States lose their feet through amputation every year. Many hundreds of thousands suffer this debilitating loss world-wide. Additionally, there are 700,000 individuals that survive a stroke each year in the United States often causing a variety of lower limb pathologies that constrain ambulation. Until recently, lower-extremity prosthetic and orthotic systems have employed predominantly passive or low-power mechanisms incapable of delivering the non-conservative positive work on each stride that the body needs to achieve an economical walking motion even on flat terrain-let alone on uneven surfaces such as stairs and steps.
It is helpful to understand the normal biomechanics associated with a gait cycle of a subject to appreciate the requirements of lower-extremity prosthetic, orthotic and exoskeleton apparatus. Specifically, the function of human ankle under sagittal plane rotation is described below for different locomotor conditions.
The mechanical characteristics of conventional passive ankle/foot prostheses (“AFPs”) like the Ossur Flex-Foot® remain essentially constant throughout the life of the device. U.S. Patent Published Application No. US 2006/0249315 (“the '315 application”) represented a significant advance over those conventional AFPs. The '315 application, the entire contents of which are hereby incorporated by reference in its entirety, recognized that performance can be improved by dividing the walking cycle into five phases, and by optimizing the mechanical characteristics of the device independently for each of those five phases.
FIG. 1A. is a schematic illustration of the different phases of a subject's gait cycle over level ground. The gait cycle is typically defined as beginning with the heel strike of one foot and ending at the next heel strike of the same foot. The gait cycle is broken down into two phases: the stance phase (about 60% of the gait cycle) and the subsequent swing phase (about 40% of the gait cycle). The swing phase represents the portion of the gait cycle when the foot is off the ground. The stance phase begins at heel-strike when the heel touches the floor and ends at toe-off when the same foot rises from the ground surface. The stance phase is separated into three sub-phases: Controlled Plantarflexion (CP), Controlled Dorsiflexion (CD), and Powered Plantarflexion (PP).
CP begins at heel-strike illustrated at 102 and ends at foot-flat at 106. CP describes the process by which the heel and forefoot initially make contact with the ground. Researchers have shown that CP ankle joint behavior is consistent with a linear spring response where joint torque is proportional to the displacement of the joint in relation to an equilibrium position of the joint position. The spring behavior is, however, variable; joint stiffness is continuously modulated by the body from step to step within the three sub-phases of stance and late swing state.
After the CP period, the CD phase continues until the ankle reaches a state of maximum dorsiflexion and begins powered plantarflexion PP as illustrated at 110. Ankle torque versus position during the CD period is described as a nonlinear spring where stiffness increases with increasing ankle position. The ankle stores the elastic energy during CD which is necessary to propel the body upwards and forwards during the PP phase.
The PP phase begins after CD and ends at the instant of toe-off illustrated at 114. During PP, the ankle applies torque in accordance with a reflex response that catapults the body upward and forward. The catapult energy is then released along with the spring energy stored during the CD phase to achieve the high plantarflexion power during late stance. This catapult behavior is necessary because the work generated during PP is more than the negative work absorbed during the CP and CD phases for moderate to fast walking speeds. The foot is lifted off the ground during the swing phase, from toe-off at 114 until the next heel strike at 118.
Because the kinematic and kinetic patterns at the ankle during stair ascent/descent are different from that of level-ground walking, a separate description of the ankle-foot biomechanics is presented in FIGS. 1B and 1C. FIG. 1B shows the human ankle biomechanics during stair ascent. The first phase of stair ascent is called Controlled Dorsiflexion 1 (CD 1), which begins with foot strike in a dorsiflexed position seen at 130 and continues to dorsiflex until the heel contacts the step surface at 132. In this phase, the ankle can be modeled as a linear spring. The second phase is Powered Plantar flexion 1 (PP 1), which begins at the instant of foot flat (when the ankle reaches its maximum dorsiflexion at 132) and ends when dorsiflexion begins once again at 134. The human ankle behaves as a torque actuator to provide extra energy to support the body weight.
The third phase is Controlled Dorsiflexion 2 (CD 2), in which the ankle dorsiflexes until heel-off at 136. For the CD 2 phase, the ankle can be modeled as a linear spring. The fourth and final phase is Powered Plantar flexion 2 (PP 2) which begins at heel-off 136 and continues as the foot pushes off the step, acting as a torque actuator in parallel with the CD 2 spring to propel the body upwards and forwards, and ends when the toe leaves the surface at 138 to begin the swing phase that ends at 140.
FIG. 1C shows the human ankle-foot biomechanics for stair descent. The stance phase of stair descent is divided into three sub-phases: Controlled Dorsiflexion 1 (CD 1), Controlled Dorsiflexion 2 (CD2), and Powered Plantar flexion (PP). CD1 begins at foot strike illustrated at 150 and ends at foot-flat 152. In this phase, the human ankle can be modeled as a variable damper. In CD2, the ankle continues to dorsiflex forward until it reaches a maximum dorsiflexion posture seen at 154. Here the ankle acts as a linear spring, storing energy throughout CD2. During PP, which begins at 154, the ankle plantar flexes until the foot lifts from the step at 156. In this final PP phase, the ankle releases stored CD2 energy, propelling the body upwards and forwards. After toe-off at 156, the foot is positioned controlled through the swing phase until the next foot strike at 158.
For stair ascent depicted in FIG. 1B, the human ankle-foot can be effectively modeled using a combination of an actuator and a variable stiffness mechanism. However, for stair descent, depicted in FIG. 1C, a variable damper needs also to be included for modeling the ankle-foot complex; the power absorbed by the human ankle is much greater during stair descent than the power released during stair ascent. Hence, it is reasonable to model the ankle as a combination of a variable-damper and spring for stair descent.
Conventional passive prosthetic, orthotic and exoskeleton apparatus do not adequately reproduce the biomechanics of a gait cycle. They are not biomimetic because they do not actively modulate impedance and do not apply the reflexive torque response; neither on level ground, ascending or descending stairs or ramps, or changing terrain conditions. A need therefore exists for improved lower-extremity prosthetic, orthotic and exoskeleton apparatus, components thereof, and methods for controlling the same.